Ultrasound imaging relies on the transmission and reception of acoustic waves. These acoustic waves are transduced from and to electrical analogous signals using electro-acoustic transducers. In diagnostic imaging, an array of closely spaced transducer elements is used. Frequently, the spacing of the elements is approximately one half or one whole wavelength according to whether or not beam steering is being used. In any event, transducer elements are tightly packed making fabrication difficult.
Ultrasound diagnostic imaging is a preferred modality for many radiology and cardiology applications. It is portable, non-invasive, involves no radiation, and can be made for relatively low cost. However, the vast majority of current ultrasound imaging systems today form 2D cross-sectional views (referred to as “B-Scans”—a 2D image plane perpendicular to the skin/transducer surface). In several applications, the B-Scan is not the preferred image format. In fact, for some applications it is preferable to use a “C-Scan”—a 2D image plane parallel to the skin/transducer surface. It is generally necessary to use a 2D transducer array and associated 2D matrix of processing channels to form a C-Scan image in real-time. Thus, the cost and complexity of the transducer array, array element signal connections and per-channel electronics and signal processing circuitry are also significant concerns.
One solution to the cost and complexity problem involves using a single common driver circuit to drive all N2 channels (where N is the number of channels along each of the two dimensions of the array) for the transmit mode of operation and using discrete channel processing only in the receive mode of operation. This approach is particularly attractive since it can be configured to reduce the number of high voltage (and therefore large, costly and bulky) processing channels to one.
It is also preferable to use as low-cost electronic circuitry as possible—typically commercial CMOS. However, commercial CMOS circuitry is designed for low voltage operation and is easily damaged by high voltages similar to those frequently used for the transmit operation (30V-150V).
Therefore, in the typical ultrasonic transducer configuration, isolation must be provided between the transmit and receive circuitry to protect the receive circuitry from the high voltages used during transmit. In systems using a shared transmit driver, the low voltage connections (i.e., receive signal paths) may be temporarily connected to ground thus protecting the channels from high potentials, while the other side of the array that is normally the shared ground during receive is pulsed with a high voltage to generate a plane wave output from all transducer elements. This type of isolation has typically been provided through the use of multiplexer switches to ground or even disconnect the receive circuitry from the transducer elements during transmission.
In some prior art systems the isolation of the transmit and receive circuits is provided by shunt circuitry. That is, during the transmission of a pulse, the receive side of the transducer may be effectively grounded using shunt circuitry, such as diodes acting as a signal clamp to divert excess signal voltages caused by the signal transmission. During the receive operation, the transmit side of the electrode is grounded to provide a receive signal reference.
FIG. 1(A) shows a prior art transducer array. Only two elements are shown for clarity. In the case of a 2D array there may be 32-64 elements along each of two dimensions. The elements in FIG. 1A are show in isolation. In addition to the elements as shown, a common backing medium is located on the bottom surface of each element. The top surface of each transducer has a protective layer (e.g. RTV silicone or polyurethane) between the transducer and the load media (i.e. patient tissue under examination.) Notice that there is one active layer (PZT) in each transducer element, while the other layer is present to provide acoustic coupling. In the case of FIG. 1A, separate ground connections are made to all PZT elements. Practically speaking, these ground connections are connected together either at the array or at some other location.
FIG. 1B is a slightly modified version of FIG. 1A. In this case, a shared ground electrode covers all the top surface of all elements. Again, note that there is only one active layer (PZT) in each transducer element, while the other layer is present to provide acoustic coupling.
The transducers of FIGS. 1A and 1B use the same transducer elements for both transmit and receive, and thus each element has two electrodes. One disadvantage of the prior art arrays is that it is not practical to simultaneously optimally match the electrical impedance of the transmit and receive circuitry in this configuration. Specifically, all the transducer channels are operated electrically in parallel in the transmit mode but in an individual electrically isolated manner in the receive mode. Because the difference in transducer electrical impedance between transmit and receive is N^2 (e.g. 1024 or 4096), the electrical matching in both transmit and receive becomes extremely problematic. Imperfect transducer matching results in wasted electrical potential drops in the transmit/receive circuitry and this results in overall reduction in signal to noise ratio (SNR) and hence reduces either potential imaging penetration and/or imaging resolution (since a lower frequency may be required to mitigate signal attenuation/loss).
The transducers of FIG. 1C are multi-layered devices. One disadvantage of these types of transducers is that they require a large number of interconnections to the individual elements.
Thus, an improved transducer array is provided that overcomes some of the disadvantages of the prior, and provides additional benefits, as described herein.